Spinal cord stimulator system

ABSTRACT

Spinal cord stimulation (SCS) system having a recharging system with self-alignment, a system for mapping current fields using a completely wireless system, multiple independent electrode stimulation outsource, and IPG control through software on Smartphone/mobile device and tablet hardware during trial and permanent implants. SCS system can include multiple electrodes, multiple, independently programmable, stimulation channels within an implantable pulse generator (IPG) providing concurrent, but unique stimulation fields. SCS system can include a replenishable power source, rechargeable using transcutaneous power transmissions between antenna coil pairs. An external charger unit, having its own rechargeable battery, can charge the IPG replenishable power source. A real-time clock can provide an auto-run schedule for daily stimulation. A bi-directional telemetry link informs the patient or clinician the status of the system, including the state of charge of the IPG battery. Other processing circuitry in current IPG allows electrode impedance measurements to be made.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation patent application of U.S. patentapplication Ser. No. 14/213,430 filed on Mar. 14, 2014, which is anon-provisional application that claims priority to provisionalapplication No. 61/792,654 filed on Mar. 15, 2013, all of which areincorporated in their entirety herein.

TECHNICAL FIELD

This disclosure relates to stimulators using electrical pulses in amedical context, and more particularly, applying electrical pulsestimulators to the spinal cord to control pain.

BACKGROUND

A Spinal Cord Stimulator (SCS) is used to exert pulsed electricalsignals to the spinal cord to control chronic pain. Spinal cordstimulation, in its simplest form, comprises stimulating electrodesimplanted in the epidural space, an electrical pulse generator implantedin the lower abdominal area or gluteal region, conducting wiresconnecting the electrodes to the electrical pulse generator, anelectrical pulse generator remote control, and an electrical pulsegenerator charger. Spinal cord stimulation has notable analgesicproperties and, at the present, is used mostly in the treatment offailed back surgery syndrome, complex regional pain syndrome andrefractory pain due to ischemia.

Electrotherapy of pain by neurostimulation began shortly after Melzackand Wall proposed the gate control theory in 1965. This theory proposedthat nerves carrying painful peripheral stimuli and nerves carryingtouch and vibratory sensation both terminate in the dorsal horn (thegate) of the spinal cord. It was hypothesized that input to the dorsalhorn of the spinal cord could be manipulated to “close the gate” to thenerves. As an application of the gate control theory, Shealy et al.implanted the first spinal cord stimulator device directly on the dorsalcolumn for the treatment of chronic pain in 1971.

Spinal cord stimulation does not eliminate pain. The electrical impulsesfrom the stimulator override the pain messages so that the patient doesnot feel the pain intensely. In essence, the stimulator masks the pain.A trial implantation is performed before implanting the permanentstimulator. The physician first implants a trial stimulator through theskin (percutaneously) to perform stimulations as a trial run. Because apercutaneous trial stimulator tends to move from its original location,it is considered temporary. If the trial is successful, the physiciancan then implant a permanent stimulator. The permanent stimulator isimplanted under the skin of the abdomen with the leads inserted underthe skin and subcutaneously fed to and inserted into the spinal canal.This placement of the stimulator in the abdomen is a more stable,effective location. The leads, which consist of an array of electrodes,can be percutaneous type or paddle type. Percutaneous electrodes areeasier to insert in comparison with paddle type, which are inserted viaincision over spinal cord and laminectomy.

There are a number of the problems that exist in currently available SCSsystems that limit the full benefits of dorsal column stimulation froman effectiveness and patient user friendly perspective. One problem isthat current SCS systems are limited to only 16 electrodes with amaximum of 16 independent current sources. Another problem is thatcurrent SCS systems have complicated trialing methods that involvemultiple gadgets and hardware. Another problem is that patients mustcarry an independent remote control in order to control the IPG in theirdaily lives.

SUMMARY

Disclosed are the following features included within a spinal cordstimulation system: (1) a recharging system with self alignment, (2) asystem for mapping current fields using a completely wireless system,(3) multiple independent electrode stimulation outsource, and (4) IPGcontrol, during trial and permanent implants, through software ongeneric Smartphone/mobile device and tablet hardware. The SCS system caninclude multiple electrodes, and multiple, independently programmable,stimulation channels within an implantable pulse generator (IPG) wherethe channels can provide concurrent, but unique stimulation fields,permitting virtual electrodes to be realized. The SCS system can includea replenishable power source (e.g., rechargeable battery) that may berecharged using transcutaneous power transmissions between antenna coilpairs. An external charger unit, having a rechargeable battery can beused to charge the IPG replenishable power source. A real-time clock canprovide an auto-run schedule for daily stimulation. An includedbi-directional telemetry link in the system can inform the patient orclinician of the status of the system, including the state of charge ofthe IPG battery. Other processing circuitry in the IPG allows electrodeimpedance measurements to be made. Circuitry provided in the externalbattery charger can provide alignment detection for the coil pairs. ASCS system, as described herein, may be for use during the trial periodand the permanent implantation.

The SCS system, as disclosed herein, is superior to existing SCS systemsin that the SCS system, as disclosed herein, can provide a stimulus to aselected pair or group of a multiplicity of electrodes, e.g., 32electrodes, grouped into multiple channels, e.g., 6 channels. In anembodiment, each electrode is able to produce a programmable constantoutput current of at least 12 mA over a range of output voltages thatmay go as high as 16 volts. In another embodiment, the implant portionof the SCS system includes a rechargeable power source, e.g., one ormore rechargeable batteries. The SCS system described herein requiresonly an occasional recharge, has an implanted portion smaller thanexisting implant systems, has a self aligning feature to guide thepatient in aligning the charger over the implanted IPG for the mostefficient power recharge, has a life of at least 10 years at typicalsettings, offers a simple connection scheme for detachably connecting alead system thereto, and is extremely reliable.

In an embodiment, each of the electrodes included within the stimuluschannels can deliver up to 12.7 mA of current over the entire range ofoutput voltages, and can be combined with other electrodes to delivercurrent up to a maximum of 20 mA. Additionally, the SCS system providesthe ability to stimulate simultaneously on all available electrodes inthe SCS system. That is, in operation, each electrode can be groupedwith at least one additional electrode to form one channel. The SCSsystem allows the activation of electrodes to at least 10 channels. Inone embodiment, such grouping is achieved by a low impedance switchingmatrix that allows any electrode contact or the system case (which maybe used as a common, or indifferent, electrode) to be connected to anyother electrode. In another embodiment, programmable output currentDAC's (digital-to-analog converters) are connected to each electrodenode, so that, when enabled, any electrode node can be grouped with anyother electrode node that is enabled at the same time, therebyeliminating the need for a low impedance switching matrix. Thisadvantageous feature allows the clinician to provide unique electricalstimulation fields for each current channel, heretofore unavailable withother “multichannel” stimulation systems (which “multi-channel”stimulation systems are really multiplexed single channel stimulationsystems).

Moreover, this feature, combined with multicontact electrodes arrangedin two or three dimensional arrays, allows “virtual electrodes” to berealized, where a “virtual electrode” comprises an electrode thatappears to be at a certain physical location, but in actuality is notphysically located at the certain physical location. Rather, the“virtual electrode” results from the vector combination of electricalfields from two or more electrodes that are activated simultaneously.

In embodiments, the SCS system includes an implantable pulse generator(IPG) powered by a rechargeable internal battery, e.g., a rechargeablelithium ion battery providing an output voltage that varies from about4.1 volts, when fully charged, to about 3.5 volts.

Embodiments are comprised of components previously not provided on SCSsystems. The components are comprised of a number of differentsub-components, as described herein. The SCS system can be comprised ofan permanent implantable IPG, an implantable trial IPG, a wirelessdongle, an IPG charger, clinical programmer software, patient programmersoftware, leads (percutaneous and paddle), lead anchors, lead splitters,lead extensions, and accessories.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts various components that can be included in a spinal cordstimulation system, according to an embodiment, during trial andpermanent implantation.

FIG. 2 depicts an exploded view of an implantable pulse generator (IPG)assembly, according to an embodiment and FIG. 2B illustrates thespecific components included in the IPG and charger.

FIG. 3 depicts a feedthrough assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 4 depicts a lead contact system of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 5 depicts a lead contact assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 6 depicts a head unit assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 7 depicts an RF antenna of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 8 depicts a percutaneous lead, according to an embodiment.

FIG. 9 depicts a paddle lead, according to an embodiment.

FIG. 10 depicts a lead extension, according to an embodiment.

FIG. 11 depicts a lead splitter, according to an embodiment.

FIG. 12 depicts a sleeve anchor, according to an embodiment.

FIG. 13 depicts a mechanical locking anchor, according to an embodiment.

FIG. 14 illustrates communication via a wireless dongle with atablet/clinician programmer and smartphone/mobile/patient programmerduring trial and/or permanent implantation, according to an embodiment.

FIG. 15 depicts a Tuohy needle, according to an embodiment.

FIG. 16 depicts a stylet, according to an embodiment.

FIG. 17 depicts a passing elevator, according to an embodiment.

FIG. 18 depicts a tunneling tool, according to an embodiment.

FIG. 19 depicts a torque wrench, according to an embodiment.

DETAILED DESCRIPTION

Implantable Pulse Generator (IPG)

FIG. 1 illustrates various components that can be included in a SCSsystem for the trial and the permanent installation periods. The spinalcord stimulator (SCS) 100 is an implantable device used to deliverelectrical pulse therapy to the spinal cord in order to treat chronicpain. The implantable components of the system consist of an ImplantablePulse Generator (IPG) 102 and a multitude of stimulation electrodes 130.The IPG 102 is implanted subcutaneously, no more than 30 mm deep in anarea that is comfortable for the patient while the stimulationelectrodes 130 are implanted directly in the epidural space. Theelectrodes 130 are wired to the IPG 102 via leads 140, 141 which keepthe stimulation pulses isolated from each other in order to deliver thecorrect therapy to each individual electrode 130.

The therapy delivered consists of electrical pulses with controlledcurrent amplitude ranging from +12.7 to −12.7 mA (current range 0-25.4mA). These pulses can be programmed in both length and frequency from 10μS to 2000 μS and 0.5 Hz to 1200 Hz. At any given moment, the sum of thecurrents sourced from the anodic electrodes must equal the sum of thecurrents sunk by the cathodic electrodes. In addition, each individualpulse is bi-phasic, meaning that once the initial pulse finishes anotherpulse of opposite amplitude is generated after a set holdoff period. Theelectrodes 130 may be grouped into stimulation sets in order to deliverthe pulses over a wider area or to target specific areas, but the sum ofthe currents being sourced at any one given time may not exceed 20 mA. Auser can also program different stimulation sets (up to eight) withdifferent parameters in order to target different areas with differenttherapies.

FIG. 2 depicts an exploded view of an IPG 102. The IPG 102 consists oftwo major active components 104, 106, a battery 108, antenna 110, somesupport circuitry, and a multitude of output capacitors 112. The firstof the major active components is the microcontroller 104. It isresponsible for receiving, decoding, and execution both commands andrequests from the external remote. If necessary it passes these commandsor requests onto the second major component, the ASIC 106. The ASIC 106receives the digital data from the microcontroller 104 and performs theentire signal processing to generate the signals necessary forstimulation. These signals are then passed onto the stimulationelectrodes 130 in the epidural space.

The ASIC 106 is comprised of a digital section and an analog section.The digital section is divided into multiple sections including; TimingGenerators, Arbitration Control, Pulse Burst Conditioner, and ElectrodeLogic. The analog section receives the incoming pulses from the digitalsection and amplifies them in order to deliver the correct therapy.There are also a multitude of digital register memory elements that eachsection utilizes, both digital and analog.

The digital elements in the ASIC 106 are all made up of standard subsetsof digital logic including logic gates, timers, counters, registers,comparators, flip-flips, and decoders. These elements are ideal forprocessing the stimulation pulses as all of them can function extremelyfast—orders of magnitudes faster than the required pulse width. Theelements all function at one single voltage, usually 5.0, 3.3, 2.5, or1.8 volts.

The timing generators are the base of each of the stimulation sets. Itgenerates the actual rising and falling edge triggers for each phase ofthe bi-phasic pulse. It accomplishes this by taking the incoming clockthat is fed from the microcontroller 104 and feeding it into a counter.For the purpose of this discussion, assume the counter simply countsthese rising clock edges infinitely. The output of the counter is fedinto six different comparators. The comparators other input is connectedto specific registers that are programmed by the microcontroller 104.When the count equals the value stored in the register, the comparatorasserts a positive signal.

The first comparator is connected to the SET signal of a SR flip flop.The SR flip flop stays positive until the RESET signal is asserted,which the second comparator is connected to. The output of the SR flipflop is the first phase of the bi-phasic pulse. Its rising & fallingedges are values stored in the registers and programmed by themicrocontroller 104. The third and fourth comparators & registers workin exactly the same way to produce the second phase of the bi-phasicpulse using the second SR flip flop.

The fifth comparator is connected the RESET of the final SR-Flip flop inthe timing generator. This flip flop is SET by the first comparator,which is the rising edge of the first pulse. The RESET is then triggeredby the value the microprocessor programmed into the register connectedto the comparator. This allows for a ‘holdoff’ period after the fallingedge of the second pulse. The output of this third SR flip flop can bethought of as an envelope of the biphasic pulses indicating when thisparticular timing generator is active.

The final comparator of the system is once again connected to a registerthat stores the frequency values from the microprocessor. Essentiallywhen the count reaches this value it triggers the comparator which isfed back to the counter to reset it to zero and beginning the entirepulse generation cycle again. The ASIC 106 may contain many of thesetiming generators as each can control anywhere from two to all of theelectrodes 130 connected to the IPG 102 at a time. However, when thereis more than one timing generator and multiple channels have beenactively programmed then there needs to be a mechanism for suppressing asecond channel from turning on when another is already active.

The next circuit block contained in the IPG 102 is the arbitrator. Thearbitrator functions by looking at each of the timing generators'envelope signals and makes sure only one can be active at a time. If asecond tries to activate then the arbitrator suppresses that signal.

The arbitrator accomplishes this by bringing each of the channelenvelope signals into a rising edge detection circuit. Once one istriggered it is fed into the SET pin of an SR flip flop. The output ofthis SR-flip flop is fed into all of the other rising edge detectors inorder to suppress them from triggering. The channel envelope signal isalso fed into a falling-edge detector which is then fed into the RESETof the same SR flip flop. The output of the SR flip flops are thenconnected to switches whose outputs are all tied together that turnon/off that channels particular biphasic pulse train. Therefore, theoutput of this circuit element is a single bi-phasic pulse train and asignal designating which timing generator that particular pulse train issourced from. Essentially, the circuit looks for a channel to go active.Once it finds one it suppresses all others until that channel becomesinactive.

The next section of the circuit works very similarly to the timinggenerators to create a high speed burst pulse train that is thencombined with the stimulation pulse train to create a bursted bi-phasicpulse train if desired.

It accomplishes this by taking the incoming clock that is fed from themicrocontroller 104 and feeding it into a counter. The counter can countthese rising clock edges infinitely. The counter is only active during asingle phase of the bi-phasic signal and begins counting as soon as therising edge is detected. The output of the counter is fed into acomparator, along with a microcontroller-programmed register, whoseoutput is connected to the reset pin on the counter.

Therefore, this counter will simply count to a programmed value andreset. This programmed value is the burst frequency.

The output of the comparator is then fed into an edge detection circuitand then a flip flop that combines it with the actual stimulation pulsetrain to create a single phase bursted stimulation pulse. The entirecircuit is duplicated for the second phase of the signal resulting inthe desired bursted bi-phasic pulse train. The stimulation signal is nowhanded over to the electrode logic stage.

The electrode logic conditions and directs the bi-phasic signals to theanalog section of the ASIC 106. At this point, the bi-phasic signalscontain all of the pertinent timing information, but none of therequired amplitude information. The incoming signals include thebi-phasic pulse train and another signal designating which timinggenerator the current active train came from. Each electrode logic cellhas a register for each timing generator that stores this particularelectrode's 130 amplitude values for that timing generator. Theelectrode logic cell uses the designation signal to determine whichregister to pull the amplitude values from, e.g. if the third timinggenerator is passed through the arbitration circuit then the electrodelogic would read the value from the third register.

Once the value is pulled from the register, it goes through a series oflogic gates. The gates first determine that the electrode 130 should beactive. If not, no further action is taken and the analog section of theelectrode output is not activated, thereby saving precious battery 108power. Next, a determination is made if the particular electrode 130 isan anode or cathode. If the electrode is deemed to be an anode, theelectrode logic passes the amplitude information and the biphasic signalto the positive current (digital to analog converter) DAC in the analogsection of the ASIC 106. If the electrode is deemed to be a cathode, theelectrode logic passes the amplitude information and the biphasic signalto the negative current DAC in the analog section of the ASIC 106. Theelectrode logic circuit must make these decisions for each phase of thebi-phasic signal as every electrode 130 will switch between being ananode and a cathode.

The analog elements in the ASIC 106 are uniquely designed in order toproduce the desired signals. The basis of analog IC design is the fieldeffect transistor (FET) and the type of high current multiple outputdesign required in SCS 100 means that the bulk of the silicon in theASIC 106 will be dedicated to the analog section.

The signals from the electrode output are fed into each current DAC whenthat specific electrode 130 should be activated. Each electrode 130 hasa positive and a negative current DAC, triggered by the electrode logicand both are never active at the same time. The job of each current DACis, when activated, to take the digital value representing a stimulationcurrent amplitude and produce an analog representation of this value tobe fed into the output stage. This circuit forms half of the barrierbetween the digital and analog sections of the ASIC 106.

The digital section of the ASIC 106 is built upon a technology that onlyallows small voltages to exist. In moving to the analog section, theoutput of the current DAC (which is a low level analog signal) must beamplified to a higher voltage for use in the analog section. The circuitthat performs this task is called a power level shifter. Because thiscircuit is built upon two different manufacturing technologies andrequires high precision analog circuits built upon a digital base, itcan be difficult to implement.

Once the voltages have been converted for usage in the analog portion ofthe ASIC 106 the voltages are passed on to the output current stages.There are two current sources per electrode output. One will source apositive current and one will sink a negative current, but both willnever be active simultaneously. The current sources themselves are madeup of analog elements similar to a Howland current source. There is aninput stage, and an amplification stage with feedback through a sensingcomponent to maintain the constant current. The input stage takes theanalog voltage values from the power level shifter and produces anoutput pulse designated for the amplifier. The amplifier then createsthe pulses of varying voltages but constant current flow. The sourcesare capable of sourcing or sinking up to 12.7 mA at 0.1 mA resolutioninto a load of up to 1.2 k Ohms. This translates into range of 15 volts,which will vary depending on the load in order to keep the currentconstant.

The microcontroller 104 to ASIC 106 interface is designed to be assimple as possible with minimal bus ‘chatter’ in order to save battery108 life. The ASIC 106 can be a collection of registers programmed via astandard I²C or SPI bus. Since the ASIC 106 is handling all the powermanagement, there will also be a power good (PG) line between the twochips in order to let the microcontroller 104 know when it is safe topower up. The ASIC 106 will also need to use a pin on themicrocontroller 104 in order to generate a hardware interrupt in caseanything goes awry in the ASIC 106. The final connection is the timebase for all of the stimulation circuitry. The ASIC 106 will require twoclocks, one for its internal digital circuitry which will be feddirectly from the microcontroller 104 clock output, and one to base allstimulation off of which will need to be synthesized by themicrocontroller 104 and fed to the ASIC 106. All commands and requeststo the ASIC 106 will be made over the I²C or SPI bus and will involvesimply reading a register address or writing to a register. Even whenthe ASIC 106 generates a hardware interrupt, it will be theresponsibility of the microcontroller 104 to poll the ASIC 106 anddetermine the cause of the interrupt.

The wireless interface is based upon the FCCs MedRadio standardoperating in the 402-405 Mhz range utilizing up to 10 channels fortelemetry. The protocol implemented is chosen to minimize transmissionand maximize battery 108 life. All processing will take place on theuser remote/programmer and the only data transmitted is exactly whatwill be used in the microcontroller 104 to ASIC 106 bus. That is, all ofthe wireless packets will contain necessary overhead information alongwith only a register address, data to store in the register, and acommand byte instructing the microcontroller 104 what to do with thedata. The overhead section of the wireless protocol will containsynchronization bits, start bytes, an address which is synchronized withthe IPG's 102 serial number, and a CRC byte to assure propertransmission. The packet length is kept as small as possible in order tomaintain battery 108 life. Since the IPG 102 cannot listen for packetsall the time due to battery 108 life, it cycles on for a duty cycle ofless than 0.05% of the time. This time value can be kept small as longas the data packets are also small. The user commands needed to run thesystem are executed by the entire system using flows.

The IPG 102 uses an implantable grade Li ion battery 108 with 215 mAHrwith zero volt technology. The voltage of the battery 108 at fullcapacity is 4.1 V and it supplies current only until it is drained up to3.3 V which is considered as 100% discharged. The remaining capacity ofthe battery 108 can be estimated at any time by measuring the voltageacross the terminals. The maximum charge rate is 107.5 mA. A ConstantCurrent, Constant Voltage (CCCV) type of regulation can be applied forfaster charging of the battery 108.

The internal secondary coil 109 is made up of 30 turns of 30 AWG coppermagnet wires. The ID, OD, and the thickness of the coil are 30, 32, and2 mm, respectively. Inductance L2 is measured to be 58 uH, a 80 nFcapacitor is connected to it to make a series resonance tank at 74 kHzfrequency. In the art of induction charging, two types of rectifiers areconsidered to convert the induced AC into usable DC, either a bridgefull wave rectifier or a voltage doubler full wave rectifier. To obtaina higher voltage, the voltage double full wave rectifier is used in thisapplication. The rectifier is built with high speed Schottky diodes toimprove its function at high frequencies of the order 100 kH. A Zenerdiode and also a 5V voltage regulator are used for regulation. Thiscircuit will be able to induce AC voltage, rectify to DC, regulate to 5Vand supply 100 mA current to power management IC that charges theinternal battery 108 by CCCV regulation.

The regulated 5V 100 mA output from the resonance tank is fed to, forexample, a Power Management Integrated Circuit (PMIC) MCP73843. Thisparticular chip was specially designed by Microchip to charge a Li ionbattery 108 to 4.1 V by CCCV regulation. The fast charge current can beregulated by changing a resistor; it is set to threshold current of 96mA in the example circuit. The chip charges the battery 108 to 4.1V aslong as the current received is more than 96 mA. However, if the supplycurrent drops below 96 mA, it stops to charge the battery 108 until thesupply is higher than 96 again. For various practical reasons, if thedistance between the coils increases, the internal secondary coil 109receives lesser current than the regulated value, and instead ofcharging the battery 108 slowly, it pauses the charging completely untilit receives more than 96 mA. It is understood to those with skill in theart that other power management chips can be used and the powermanagement chip is not limited to the PMIC MCP738432 chip.

All the functions of the IPG 102 are controlled from outside using ahand held remote controller specially designed for this device. Alongwith the remote control, an additional control is desirable to operatethe IPG 102 if the remote control was lost or damaged. For this purposea Hall effect based magnet switch was incorporated to either turn ON orturn OFF the IPG 102 using an external piece of magnet. Magnet switchacts as a master control for the IPG 102 to turn on or off. A south poleof sufficient strength turns the output on and a north pole ofsufficient strength turns the output off. The output is latched so thatthe switch continues to hold the state even after the magnet is removedfrom its vicinity.

The IPG 102 is an active medical implant that generates an electricalsignal that stimulates the spinal cord. The signal is carried through astimulation lead 140 that plugs directly into the IPG 102. The IPG 102recharges wirelessly through an induction coil 109, and communicates viaRF radio antenna 110 to change stimulation parameters. The IPG 102 isimplanted up to 3 cm below the surface of the skin and can be fixed tothe fascia by passing two sutures through holes in the epoxy header 114.The leads 140 are electrically connected to the IPG 102 through a leadcontact system 116, a cylindrical spring-based contact system withinter-contact silicone seals. The leads 140 are secured to the IPG 102with a set screw that actuates within locking housing 118. Set screwcompression on the lead's 140 fixation contact can be governed by adisposable torque wrench. The wireless recharging is achieved byaligning the exterior induction coil on the charger with the internalinduction coil 109 within the IPG 102. The RF antenna within theremote's dongle 200 communicates with the RF antenna 110 in the IPG's102 epoxy header 114. FIG. 2 illustrates an exploded view of the IPG 102assembly.

The IPG 102 is an assembly of a hermetic titanium (6Al-4V) casing 120which houses the battery 108, circuitry 104, 106, and charging coil 109.The IPG 102 further includes an epoxy header 114 (see FIG. 6), whichhouses the lead contact assembly 116, housing 118, and RF antenna 110(see FIGS. 6 and 7). The internal electronics are connected to thecomponents within the epoxy head through a hermetic feedthrough 122, asshown in FIG. 3. The feedthrough 122 is a titanium (6Al-4V) flange withan alumina window and gold trimming. Within the alumina window arethirty-four platinum-iridium (90-10) pins that interface internally witha direct solder to the circuit board, and externally with a series ofplatinum iridium wires laser-welded to the antenna 110 and lead contacts126. The IPG 102 interfaces with 32 lead contacts 126, which arearranged in four rows of eight contacts 126. Thirty two of thefeedthrough's 122 pins 124 interface with the contacts 126, while twointerface with the antenna 110, one to the ground plane and one to theantenna 110 feed.

FIGS. 4 and 5 depict a lead contact system 115 and assembly 116,respectively. The lead contacts 126 consist of an MP35N housing 128 witha platinum-iridium 90-10 spring 129. Each contact 126 is separated by asilicone seal 127. At the proximal end of each stack of 8 contacts 126is a titanium (6Al-4V) cap 125 which acts as a stop for the lead 140. Atthe distal end is a titanium (6Al-4V) set screw and housing 118 for leadfixation. At the lead entrance point is a silicone tube 123 whichprovides strain relief as the lead 140 exits the head unit 114, andabove the set screw hole 119 another silicone tube 131 with a smallinternal canal allows the torque wrench to enter but does not allow theset screw 449 to back out. In addition to the contacts 126 and antenna110, the header 114 also contains a radiopaque titanium (6Al-4V) tag 132which allows for identification of the device under fluoroscopy. Theovermold of the header 114 is Epotek 301, a two-part, biocompatibleepoxy. FIGS. 4, 5, 6, and 7 depict illustrations of lead contact system115, lead contact assembly 116, head unit assembly 114, and RF antenna110, respectively.

Internal to the titanium (6Al-4V) case 120 are the circuit board 105,battery 108, charging coil 109, and internal plastic support frame. Thecircuit board 105 can be a multi-layered FR-4 board with copper tracesand solder mask coating. Non-solder masked areas of the board can beelectroless nickel immersion gold. The implantable battery 108, allsurface mount components, ASIC 106, microcontroller 104, charging coil109, and feedthrough 122 will be soldered to the circuit board 105. Theplastic frame, made of either polycarbonate or ABS, will maintain thebattery's 108 position and provide a snug fit between the circuitry 105and case 120 to prevent movement. The charging coil 109 is a woundcoated copper.

Leads

The percutaneous stimulation leads 140, as depicted in FIG. 8, are afully implantable electrical medical accessory to be used in conjunctionwith the implantable SCS 100. The primary function of the lead is tocarry electrical signals from the IPG 102 to the target stimulation areaon the spinal cord. Percutaneous stimulation leads 140 providecircumferential stimulation. The percutaneous stimulation leads 140provide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and stimulation area. The leads 140 are surgicallyimplanted through a spinal needle, or epidural needle, and are driventhrough the spinal canal using a steering stylet that passes through thecenter of the lead 140. The leads 140 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 140. The leads 140 are secured at theproximal end with a set-screw 119 on the IPG 102 which applies radialpressure to a blank contact on the distal end of the proximal contacts.

The percutaneous stimulation leads 140 consist of a combination ofimplantable materials. Stimulation electrodes 130 at the distal end andelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy. This alloy is utilized for its bio-compatibilityand electrical conductivity. The electrodes 130 are geometricallycylindrical. The polymeric body of the lead 140 is polyurethane, chosenfor its bio-compatibility, flexibility, and high lubricity to decreasefriction while being passed through tissue. The polyurethane tubing hasa multi-lumen cross section, with one center lumen 142 and eight outerlumens 144. The center lumen 142 acts as a canal to contain the steeringstylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. The wires 146are laser welded to the contacts and electrodes 130, creating anelectrical connection between respective contacts on the proximal anddistal ends. The leads 140 employ a platinum-iridium plug 148, moldedinto the distal tip of the center lumen 142 to prevent the tip of thesteering stylet from puncturing the distal tip of the lead 140. Leads140 are available in a variety of 4 and 8 electrode 130 configurations.These leads 140 have 4 and 8 proximal contacts (+1 fixation contact),respectively. Configurations vary by electrode 130 number, electrode 130spacing, electrode 130 length, and overall lead 140 length.

The paddle stimulation leads 141, as depicted in FIG. 9, are a fullyimplantable electrical medical accessory to be used in conjunction withthe implantable SCS 100. The primary function of the paddle lead 141 isto carry electrical signals from the IPG 102 to the target stimulationarea on the spinal cord. The paddle leads 141 provide uni-directionstimulation across a 2-dimensional array of electrodes 130, allowing forgreater precision in targeting stimulation zones. The paddle stimulationleads 141 provide a robust, flexible, and bio-compatible electricconnection between the IPG 102 and stimulation area. The leads 141 aresurgically implanted through a small incision, usually in conjunctionwith a laminotomy or laminectomy, and are positioned using forceps or asimilar surgical tool. The leads 141 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 141. The leads 141 are secured at theproximal end with a set-screw on the IPG 102 which applies radialpressure to a fixation contact on the distal end of the proximalcontacts.

The paddle stimulation leads 141 consist of a combination of implantablematerials. Stimulation electrodes 130 at the distal end and electricalcontacts at the proximal end are made of a 90-10 platinum-iridium alloyutilized for its bio-compatibility and electrical conductivity. Thepolymeric body of the lead 141 is polyurethane, chosen for itsbio-compatibility, flexibility, and high lubricity to decrease frictionwhile being passed through tissue. The polyurethane tubing has amulti-lumen cross section, with one center lumen 142 and eight outerlumens 144. The center lumen 142 acts as a canal to contain the steeringstylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. At the distaltip of the paddle leads 141 is a 2-dimensional array of flat rectangularelectrodes 130 molded into a flat silicone body 149. In an embodiment,one side of the rectangular electrodes 130 is exposed, providinguni-directional stimulation. The wires 146 are laser welded to thecontacts and electrodes 130, creating an electrical connection betweenrespective contacts on the proximal and distal ends. Also molded intothe distal silicone paddle is a polyester mesh 147 adding stability tothe molded body 149 while improving aesthetics by covering wire 146routing. The number of individual 8-contact leads 141 used for eachpaddle 141 is governed by the number of electrodes 130. Electrodes 130per paddle 141 range from 8 to 32, split into between one and fourproximal lead 141 ends. Each proximal lead 141 has 8 contacts (+1fixation contact). Configurations vary by electrode 130 number,electrode 130 spacing, electrode length, and overall lead length.

The lead extensions 150, as depicted in FIG. 10, are a fully implantableelectrical medical accessory to be used in conjunction with theimplantable SCS 100 and either percutaneous 140 or paddle 141 leads. Theprimary function of the lead extension 150 is to increase the overalllength of the lead 140, 141 by carrying electrical signals from the IPG102 to the proximal end of the stimulation lead 140, 141. This extendsthe overall range of the lead 140, 141 in cases where the length of theprovided leads 140, 141 are insufficient. The lead extensions 150provide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and proximal end of the stimulation lead 140, 141.The extensions 150 may be secured mechanically to the patient usingeither an anchor or a suture passed through tissue and tied around thebody of the extension 150. Extensions 150 are secured at the proximalend with a set-screw 119 on the IPG 102 which applies radial pressure toa fixation contact on the distal end of the proximal contacts of theextension 150. The stimulation lead 140, 141 is secured to the extension150 in a similar fashion, using a set screw 152 inside the molded tip ofextension 150 to apply a radial pressure to the fixation contact at theproximal end of the stimulation lead 140, 141.

The lead extension 150 consists of a combination of implantablematerials. At the distal tip of the extension 150 is a 1×8 array ofimplantable electrical contacts 154, each consisting of MP35 housing 128and 90-10 platinum-iridium spring. A silicone seal 127 separates each ofthe housings 128. At the proximal end of the contacts is a titanium(6Al4V) cap which acts as a stop for the lead, and at the distal tip, atitanium (6Al4V) block and set screw 152 for lead fixation. Theelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy utilized for its bio-compatibility and electricalconductivity. The polymeric body 156 of the lead 150 is polyurethane,chosen for its bio-compatibility, flexibility, and high lubricity todecrease friction while being passed through tissue. The polyurethanetubing 158 has a multi-lumen cross section, with one center lumen 142and eight outer lumens 144. The center lumen 142 acts as a canal tocontain the steering stylet during implantation, while the outer lumens144 provide electrical and mechanical separation between the wires 146that carry stimulation from the proximal contacts to distal electrodes.These wires 146 are a bundle of MP35N strands with a 28% silver core.The wires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. Each leadextension 150 has 8 proximal cylindrical contacts (+1 fixation contact).

The lead splitter 160, as depicted in FIG. 11, is a fully implantableelectrical medical accessory which is used in conjunction with the SCS100 and typically a pair of 4-contact percutaneous leads 140. Theprimary function of the lead splitter 160 is to split a single lead 140of eight contacts into a pair of 4 contact leads 140. The splitter 160carries electrical signals from the IPG 102 to the proximal end of two4-contact percutaneous stimulation leads 140. This allows the surgeonaccess to more stimulation areas by increasing the number of stimulationleads 140 available. The lead splitter 160 provides a robust, flexible,and bio-compatible electrical connection between the IPG 102 andproximal ends of the stimulation leads 140. The splitters 160 may besecured mechanically to the patient using either an anchor or a suturepassed through tissue and tied around the body of the splitter 160.Splitters 160 are secured at the proximal end with a set-screw 119 onthe IPG 102 which applies radial pressure to a fixation contact on thedistal end of the proximal contacts of the splitter 160. The stimulationleads 140 are secured to the splitter 160 in a similar fashion, using apair of set screws inside the molded tip of splitter 160 to apply aradial pressure to the fixation contact at the proximal end of eachstimulation lead 140.

The lead splitter 160 consists of a combination of implantablematerials. At the distal tip of the splitter 160 is a 2×4 array ofimplantable electrical contacts 162, with each contact 162 consisting ofMP35 housing 128 and 90-10 platinum-iridium spring. A silicone seal 127separates each of the housings 128. At the proximal end of each row ofcontacts 162 is a titanium (6Al4V) cap which acts as a stop for thelead, and at the distal tip, a titanium (6Al4V) block and set screw forlead fixation. The electrical contacts at the proximal end of thesplitter 160 are made of a 90-10 platinum-iridium alloy utilized for itsbio-compatibility and electrical conductivity. The polymeric body 164 ofthe lead 160 is polyurethane, chosen for its bio-compatibility,flexibility, and high lubricity to decrease friction while being passedthrough tissue. The polyurethane tubing 166 has a multi-lumen crosssection, with one center lumen 142 and eight outer lumens 144. Thecenter lumen 142 acts as a canal to contain the steering stylet duringimplantation, while the outer lumens 144 provide electrical andmechanical separation between the wires 146 that carry stimulation fromthe proximal contacts to distal electrodes 130. These wires 146 are abundle of MP35N strands with a 28% silver core. The wires 146 areindividually coated with ethylene tetrafluoroethylene (ETFE), to providean additional non-conductive barrier. Each lead splitter 160 has 8proximal contacts (+1 fixation contact), and 2 rows of 4 contacts 162 atthe distal end.

Anchors

The lead anchor 170, as depicted in FIGS. 12 and 13, is a fullyimplantable electrical medical accessory which is used in conjunctionwith both percutaneous 140 and paddle 141 stimulation leads. The primaryfunction of the lead anchor 170 is to prevent migration of the distaltip of the lead 140, 141 by mechanically locking the lead 140, 141 tothe tissue. There are currently two types of anchors 170, a simplesleeve 171, depicted in FIG. 12, and a locking mechanism 172, depictedin FIG. 13, and each has a slightly different interface. For the simplesleeve type anchor 171, the lead 140, 141 is passed through the centerthru-hole 174 of the anchor 171, and then a suture is passed around theoutside of the anchor 171 and tightened to secure the lead 140, 141within the anchor 171. The anchor 171 can then be sutured to the fascia.The locking anchor 172 uses a set screw 176 for locking purposes, and abi-directional disposable torque wrench for locking and unlocking.Tactile and audible feedback is provided for both locking and unlocking.

Both anchors 171, 172 can be molded from implant-grade silicone, but thelocking anchor 172 uses an internal titanium assembly for locking. The3-part mechanism is made of a housing 175, a locking set screw 176, anda blocking set screw 177 to prevent the locking set screw from back out.All three components can be titanium (6Al4V). The bi-directional torquewrench can have a plastic body and stainless steel hex shaft.

Wireless Dongle

The wireless dongle 200 is the hardware connection to asmartphone/mobile 202 or tablet 204 that allows communication betweenthe trial generator 107 or IPG 102 and the smartphone/mobile device 202or tablet 204, as illustrated in FIG. 14. During the trial or permanentimplant phases, the wireless dongle 200 is connected to the tablet 204through the tablet 204 specific connection pins and the clinicianprogrammer software on the tablet 204 is used to control the stimulationparameters. The commands from the clinician programmer software aretransferred to the wireless dongle 200 which is then transferred fromthe wireless dongle 200 using RF signals to the trial generator 107 orthe IPG 102. Once the parameters on the clinician programmers have beenset, the parameters are saved on the tablet 204 and can be transferredto the patient programmer software on the smartphone/mobile device 202.The wireless dongle 200 is composed of an antenna, a microcontroller(having the same specifications as the IPG 102 and trial generator 107),and a pin connector to connect with the smartphone/mobile device 202 andthe tablet 204.

Charger

The IPG 102 has a rechargeable lithium ion battery 108 to power itsactivities. An external induction type charger 210 (FIG. 1) wirelesslyrecharges the included battery 108 inside the IPG 102. The charger 210is packaged into a housing and consists of a rechargeable battery 205, aprimary coil 203 of wire and a printed circuit board (PCB) 211containing the electronics. In operation, charger 210 produces amagnetic field and induces voltage into the secondary coil 109 in theIPG 102. The induced voltage is then rectified and used to charge thebattery 108 inside the IPG 102. To maximize the coupling between thecoils, both internal and external coils are combined with capacitors 213to make them resonate at a particular common frequency. The coil 203acting as an inductor L forms an LC resonance tank. The charger 210 usesa Class-E amplifier 209 topology to produce the alternating current inthe primary coil 203 around the resonant frequency. The charger 210features include, but are not limited to:

-   -   Charge IPG 102 wirelessly    -   Charge up to a maximum depth of 30 mm    -   Integrated alignment sensor 207 indicates alignment between the        charger and IPG 102 resulting in higher power transfer        efficiency    -   Alignment sensor 207 provides audible and visual feedback to the        user    -   Compact and Portable

A protected type of cylindrical Li ion battery is used as the charger210 battery. A Class-E power amplifier topology is a much used type ofamplifier for induction chargers, especially for implantable electronicmedical devices. Due to the Class-E power amplifier's relatively hightheoretical efficiency it is often used for devices where highefficiency power transfer is necessary. A 0.1 ohm high wattage resistoris used in series to sense the current through this circuit.

The primary coil L1 is made by 60 turns of Litz wire type 100/44-100strands of 44 AWG each. The Litz wire solves the problem of skin effectand keeps its impedance low at high frequencies. Inductance of this coilwas initially set at 181 uH, but backing it with a Ferrite plateincreases the inductance to 229.7 uH. The attached ferrite plate focusesthe produced magnetic field towards the direction of the implant. Such asetup helps the secondary coil receive more magnetic fields and aids itto induce higher power.

When the switch is ON, the resonance is at frequency

$f = \frac{1}{2\;\pi\sqrt{L\; 1C\; 2}}$When the switch is OFF, it shifts to

$f = \frac{1}{2\;\pi\sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$In a continuous operation the resonance frequency will be in the range

$\frac{1}{2\;\pi\sqrt{L\; 1C\; 2}} < f < \frac{1}{2\;\pi\sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$

To make the ON and OFF resonance frequencies closer, a relatively largervalue of C1 can be chosen by a simple criteria as follows

C1=nC2; a value of n=4 was used in the example above; in most cases3<n<10.

The voltages in these Class-E amplifiers typically go up to the order of300 VAC. Capacitors selected must be able to withstand these highvoltages, sustain high currents and still maintain low Effective SeriesResistance (ESR). Higher ESRs result in unnecessary power losses in theform of heat. The circuit is connected to the battery through aninductor which acts as a choke. The choke helps to smoothen the supplyto the circuit. The N Channel MOSFET acts as a switch in this Class-Epower amplifier. A FET with low ON resistance and with high draincurrent Id is desirable.

In summary, the circuit is able to recharge the IPG 102 battery 108 from0 to 100% in approximately two hours forty-five minutes with distancebetween the coils being 29 mm. The primary coil and the Class-Eamplifier draws DC current of 0.866 A to achieve this task. To improvethe efficiency of the circuit, a feedback closed loop control isimplemented to reduce the losses. The losses are minimum when the MOSFETis switched ON and when the voltage on its drain side is close to zero.

The controller takes the outputs from operational amplifiers, checks ifthe outputs meet the criteria, then triggers the driver to switch ON theMOSFET for the next cycle. The controller can use a delay timer, an ORgate and a 555 timer in monostable configuration to condition the signalfor driver. When the device is switched ON, the circuit will notfunction right away as there is no active feedback loop. The feedbackbecomes active when the circuit starts to function. To provide an activefeedback loop, an initial external trigger is applied to jump start thesystem.

Alignment Sensor

The efficiency of the power transfer between the external charger 210and the internal IPG 102 will be maximum only when the charger 210 andIPG 102 are properly aligned. An alignment sensor is provided to ensureproper alignment as part of the external circuit design and is based onthe principle of reflected impedance. When the external coil is broughtcloser to the internal coil, the impedance of both circuits change. Thesensing is based on measuring the reflected impedance and testingwhether it crosses the threshold. A beeper provides an audible feedbackto the patient and a LED provides visual feedback.

When the impedance of the circuit changes, the current passing throughit also changes. A high power 0.1 ohm resistor can be used in the seriesof the circuit to monitor the change in current. The voltage drop acrossthe resistor is amplified 40 times and then compared to a fixedthreshold value using an operational amplifier voltage comparator. Theoutput is fed to a timer chip which in turn activates the beeper and LEDto provide feedback to the user.

The circuit can sense the alignment up to a distance of approximately 30mm. The current fluctuation in the circuit depends on more factors thanreflected impedance alone and the circuit is sensitive to otherparameters of the circuit as well. To reduce the sensitivity related toother parameters, one option is to eliminate interference of all theother factors and improve the functionality of the reflected impedancesensor—which is very challenging to implement within the limited spaceavailable for circuitry. Another option is to use a dedicated sensorchip to measure the reflected impedance.

A second design uses sensors designed for proximity detector or metaldetectors for alignment sensing. Chips designed to detect metal bodiesby the effect of Eddy currents on the HF losses of a coil can be usedfor this application. The TDE0160 is an example of such a chip.

The external charger is designed to work at 75 to 80 kHz, whereas theproximity sensor was designed for 1 MHz. The sensor circuit is designedto be compatible with the rest of the external and is fine tuned todetect the internal IPG 102 from a distance of approximately 30 mm.

Programmer

The Clinician Programmer is an application that is installed on a tablet204. It is used by the clinician to set the stimulation parameters onthe trial generator 107 or IPG 102 during trial and permanentimplantation in the operating room. The clinician programmer is capableof saving multiple settings for multiple patients and can be used toadjust the stimulation parameters outside of the operations room. It iscapable of changing the stimulation parameters though the RF wirelessdongle 200 when the trial generator 107 or IPG 102 which has beenimplanted in the patient is within the RF range. In addition, it is alsocapable of setting or changing the stimulation parameters on the trialgenerator 107 and/or the IPG 102 through the internet when both thetablet 204 and the Patient Programmers on a smartphone/mobile device 202both have access to the internet.

The Patient Programmer is an application that is installed on asmartphone/mobile device 202. It is used by the patient to set thestimulation parameters on the trial generator 107 or IPG 102 after trialand permanent implantation outside the operating room. The clinicianprogrammer is capable of saving multiple settings for multiple patientsand can be transferred to the Patient Programmer wirelessly when theClinician Programmer tablet 204 and the Patient Programmersmartphone/mobile device 202 are within wireless range such as Bluetoothfrom each other. In the scenario where the Clinician Programmer tablet204 and the Patient Programmer smartphone/mobile device 202 are out ofwireless range from each other, the data can be transferred through theinternet where both devices 202, 204 have wireless access such as Wi-Fi.The Patient Programmer is capable of changing the stimulation parameterson the trial generator 107 or IPG 102 though the RF wireless dongle 200when the trial generator 107 or IPG implanted in the patient is withinthe RF range.

Tuohy Needle

The tuohy needle 240, as depicted in FIG. 15, is used in conjunctionwith a saline-loaded syringe for loss-of-resistance needle placement,and percutaneous stimulation leads 140, for lead 140 placement into thespinal canal. The tuohy epidural needle 240 is inserted slowly into thespinal canal using a loss-of-resistance technique to gauge needle 240depth. Once inserted to the appropriate depth, the percutaneousstimulation lead 140 is passed through the needle 240 and into thespinal canal.

The epidural needle 240 is a non-coring 14G stainless steel spinalneedle 240 and will be available in lengths of 5″ (127 mm) and 6″(152.4). The distal tip 242 of the needle 240 has a slight curve todirect the stimulation lead 140 into the spinal canal. The proximal end246 is a standard Leur-Lock connection 248.

Stylet

The stylet 250, as depicted in FIG. 16, is used to drive the tip of apercutaneous stimulation lead 140 to the desired stimulation zone byadding rigidity and steerability. The stylet 250 wire 252 passes throughthe center lumen 142 of the percutaneous lead 140 and stops at theblocking plug at the distal tip of the lead 140. The tip of the stylet250 comes with both straight and curved tips. A small handle 254 is usedat the proximal end of the stylet 250 to rotate the stylet 250 withinthe center lumen 142 to assist with driving. This handle 254 can beremoved and reattached allowing anchors 170 to pass over the lead 140while the stylet 250 is still in place. The stylet 250 wire 252 is aPTFE coated stainless steel wire and the handle 254 is plastic.

Passing Elevator

The passing elevator 260, as depicted in FIG. 17, is used prior topaddle lead 141 placement to clear out tissue in the spinal canal andhelp the surgeon size the lead to the anatomy. The passing elevator 260provides a flexible paddle-shaped tip 262 to clear the spinal canal ofobstructions. The flexible tip is attached to a surgical handle 264.

The passing elevator 260 is a one-piece disposable plastic instrumentmade of a flexible high strength material with high lubricity. Theflexibility allows the instrument to easily conform to the angle of thespinal canal and the lubricity allows the instrument to easily passthrough tissue.

Tunneling Tool

The tunneling tool 270, as depicted in FIG. 18, is used to provide asubcutaneous canal to pass stimulation leads 140 from the entrance pointinto the spinal canal to the IPG implantation site. The tunneling tool270 is a long skewer-shaped tool with a ringlet handle 272 at theproximal end 274. The tool 270 is covered by a plastic sheath 276 with atapered tip 278 which allows the tool 270 to easily pass through tissue.Once the IPG 102 implantation zone is bridge to the lead 140 entrancepoint into the spinal canal, the inner core 275 is removed, leaving thesheath 276 behind. The leads 140 can then be passed through the sheath276 to the IPG 102 implantation site. The tunneling tool 270 is oftenbent to assist in steering through the tissue.

The tunneling tool 270 is made of a 304 stainless steel core with afluorinated ethylene propylene (FEP) sheath 276. The 304 stainless steelis used for its strength and ductility during bending, and the sheath276 is used for its strength and lubricity.

Torque Wrench

The torque wrench 280, as depicted in FIG. 19, is used in conjunctionwith the IPG 102, lead extension 150 and lead splitter 160 to tightenthe internal set screw 119, which provides a radial force against thefixation contact of the stimulation leads 140, 141, preventing the leads140, 141 from detaching. The torque wrench 280 is also used to lock andunlock the anchor 170. The torque wrench 280 is a small, disposable,medical instrument that is used in every SCS 100 case. The torque wrench280 provides audible and tactile feedback to the surgeon that the lead140, 141 is secured to the IPG 102, extension 150, or splitter 160, orthat the anchor 170 is in the locked or unlocked position.

The torque wrench 280 is a 0.9 mm stainless steel hex shaft 282assembled with a plastic body 284. The wrench's 280 torque rating isbi-directional, primarily to provide feedback that the anchor 170 iseither locked or unlocked. The torque rating allows firm fixation of theset screws 119, 152 against the stimulation leads 140, 141 withoutover-tightening.

Trial Patch

The trial patch is used in conjunction with the trialing pulse generator107 to provide a clean, ergonomic protective cover of the stimulationlead 140, 141 entrance point in the spinal canal. The patch is alsointended to cover and contain the trial generator 107. The patch is alarge, adhesive bandage that is applied to the patient post-operativelyduring the trialing stage. The patch completely covers the leads 140,141 and generator 107, and fixates to the patient with anti-microbialadhesive.

The patch is a watertight, 150 mm×250 mm anti-microbial adhesive patch.The watertight patch allows patients to shower during the trialingperiod, and the anti-microbial adhesive decreases the risk of infection.The patch will be made of polyethylene, silicone, urethane, acrylate,and rayon.

Magnetic Switch

The magnetic switch is a magnet the size of a coin that, when placednear the IPG 102, can switch it on or off. The direction the magnet isfacing the IPG 102 determines if the magnetic switch is switching theIPG 102 on or off.

The invention claimed is:
 1. A method of providing electrical pulsetherapy to patient via an implantable spinal cord stimulation device,comprising the steps of: providing an implantable pulse generator with amicrocontroller and an application specific integrated circuit (ASIC);receiving one or more commands sent by an external remote at atransceiver and transmitting the one or more commands to themicrocontroller; receiving, by the ASIC, data representing the one morecommands; performing, by the ASIC, signal processing in accordance withthe one or more commands to generate one or more signals for theelectrical pulse therapy; and generating a high speed burst pulse andcombining the high speed burst pulse with a stimulation pulse togenerate a bursted bi-phasic pulse train, and wherein delivering the oneor more signals for electrical pulse therapy comprises delivering thebursted bi-phasic pulse train wherein the bi-phasic pulse train includestiming information and does not contain amplitude information,delivering the one or more signals for the electrical pulse therapy toelectrodes positioned in an epidural space of a patient.
 2. The methodof claim 1, wherein the step of providing the implantable pulsegenerator further includes a step of providing a circuit assembly housedin a casing, the circuit assembly including a plurality of outputcapacitors positioned within the implantable pulse generator.
 3. Themethod of claim 1, wherein the step of providing the implantable pulsegenerator further includes a step of providing at least one rechargeableimplantable battery which is housed in the casing and operably connectedwith the circuit assembly.
 4. The method of claim 1, further comprisingcharging the implantable pulse generator via an external chargerincluding a primary charging coil and at least one battery.
 5. Themethod of claim 4, wherein the step of providing the implantable pulsegenerator further comprises providing a circuit board operably connectedwith the primary charging coil and the at least one rechargeablebattery.
 6. The method of claim 5, further comprising recharging theimplantable pulse generator, via the at least one rechargeable battery,using transcutaneous power transmissions between the primary chargingcoil and a secondary charging coil disposed in the implantable pulsegenerator.
 7. The method of claim 1, wherein the method furthercomprises the step of providing an analog and a digital circuit.
 8. Amethod of providing electrical pulse therapy to patient via animplantable spinal cord stimulation device, comprising the steps of:providing an implantable pulse generator with a circuit assembly,wherein the circuit assembly includes a microcontroller and anapplication specific integrated circuit (ASIC); receiving one or morecommands sent by an external remote at a transceiver and transmittingthe one or more commands to the microcontroller; receiving, by the ASIC,data representing the one more commands performing, by the ASIC, signalprocessing in accordance with the one or more commands to generate oneor more signals for the electrical pulse therapy; and positioningelectrodes on a patient in an epidural space, delivering the one or moresignals for the electrical pulse therapy to the epidural space of apatient, generating a high speed burst pulse and combining the highspeed burst pulse with a stimulation pulse to generate a burstedbi-phasic pulse train, and wherein delivering the one or more signalsfor electrical pulse therapy comprises delivering the bursted bi-phasicpulse train wherein the electrical pulse therapy is delivered by one ormore electrodes disposed in the epidural space wherein the bi-phasicpulse train includes timing information and does not contain amplitudeinformation.
 9. The method of claim 8, further comprising providing aplurality of output capacitors positioned within the implantable pulsegenerator.
 10. The method of claim 8, wherein the method comprises thestep of providing at least one rechargeable implantable battery ishoused in the casing and operably connected with the circuit assembly.11. The method of claim 8, further comprising charging the implantablepulse generator via an external charger including a primary chargingcoil and at least one battery.
 12. The method of claim 11, furthercomprises providing a circuit board operably connected with the primarycharging coil and the at least one rechargeable battery.
 13. The methodof claim 12, further comprising recharging the implantable pulsegenerator, via the at least one rechargeable battery, usingtranscutaneous power transmissions between the primary charging coil anda secondary charging coil disposed in the implantable pulse generator.14. The method of claim 8, the method further comprises the step ofproviding an analog and a digital circuit.